The field of the invention is nuclear magnetic resonance imaging methods and systems. More particularly, the invention relates to the Imaging of ultrasonic fields using a magnetic resonance imaging (MRI) system.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B.sub.0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B.sub.1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, M.sub.z, may be rotated, or "tipped", into the x-y plane to produce a net transverse magnetic moment M.sub.t. A signal is emitted by the excited spins after the excitation signal B.sub.1 is terminated, this signal may be received and processed to form an image.
When utilizing these signals to produce images, magnetic field gradients (G.sub.x G.sub.y and G.sub.z) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
Nuclear magnetic resonance signals can be influenced by motion through the application of a magnetic field gradient superimposed on the static, uniform magnetic field that is used for spin polarization (Hahn E L, Detection of c-water motion by nuclear precession, J Geo-phys Respectively 1960;65:766-777). This concept has been used with proton MRI to image fluid flow (Frayne R, Steinmann D A, Ethier C R, Rutt B K, Accuracy of MR phase contrast velocity measurements for unsteady flow, J Magn. Reson. Imaging 1995;5:428-431), brain (Enzmann D R, Pelc N J, Brain motion: measurement with phase contrast MR imaging, Radiology 1992;185:653-600) and muscle (Pelc L R, Sayre J, Yun K, Castro L J, Herfkens R J, Miller D C, Pelc N J, Evaluation of myocardial motion tracking with cine-phase contrast magnetic resonance imaging, Invest Radiol 1994;29:1038-1042) tissue motion, as well as the measurement of cardiac tissue strain (Wedeen V J, Magnetic resonance imaging of myocardial kinematics, Technique to detect, localize, and quantify the strain rates of the active human myocardium, Magn Reson Med 1992;27:52-67). Several authors have used oscillating magnetic field gradients to detect oscillatory fluid flow or viscoelastic tissue motion. Specifically, acoustic oscillatory fluid flow in the rat cochlea (Denk W, Keolian R M, Ogawa S. Jelinski L W, Oscillatory flow in the cochlea visualized by a magnetic resonance imaging technique, Proc Natl Acad Sci 1993; 90:1595-1598) has been demonstrated at frequencies up to 4.6 kHz. Similarly, the viscoelastic properties of tissue have been measured by monitoring the wave velocity of slow (&lt;10 m/s), low-frequency shear waves (&lt;1.1 kHz) (Lewa C J, de Certaines J D, Viscoelastic property detection by elastic displacement NMR measurements, J Magn Reson Imaging 1996; 6:652-656; Muthupillai R, Lomas D J, Rossman P J, Greenleaf J F, Manduca A, Ehman R L, Magnetic resonance elastography by direct visualization of propagation acoustic strain waves, Science 1995; 269:1854-1857; Muthupillai R, Rossman P J, Lomas D J, Greenleaf J F, Riederer S J, Ehman R L, Magnetic resonance imaging of transverse acoustic strain waves, Magn Reson Med 1996; 36:266-274) generated by mechanical stimulation. In these cases, the oscillation frequency was limited to a few kHz, with motion amplitudes ranging from 200-1000 nanometers.
The concept of motion-sensitized MRI arises from the fact that the proton Larmor frequency is proportional to the local magnetic field. By using a magnetic field gradient, in addition to a uniform polarizing field, the resultant distribution of frequencies encodes the spatial distribution of spin density. Similarly, when spins are moving in the presence of this gradient, the phase of the spins indicates the history of spin location in their movement through a gradient.
More specifically, the phase of transverse magnetization is influenced by its location and the presence of an applied magnetic field gradient. As described in more detail below, the motion of a spin can be measured by this phase depending on the nature of the motion, and the duration, waveform and amplitude of the applied magnetic field gradient.
The application of ultrasound (US) in the medical field is widespread and growing rapidly. The success of these diverse applications is largely determined by the ability to craft specific acoustical field patterns within tissue in a controlled and predictable fashion. The ability to observe a field pattern in acoustically heterogeneous tissues is important to understand the effect of phase aberrations on spatial resolution in imaging application, and on the deposition of thermal energy in high-intensity focused ultrasound (HIFU) applications. Validation of US field patterns in homogeneous media can be predicted on the basis of classical diffraction theory (Hunt J W, Arditi M, Foster F S, Ultrasound transducers for pulse-echo medical imaging, IEEE Trans Biomed Eng 1983; BME-30:453-481) and verified with invasive sensors (Schafer M E, Lewin P A, Computerized system for measuring the acoustic output from diagnostic ultrasound equipment, IEEE Trans Ultrason Ferroelectr Freq Control 1988;35:102-109; Fry W S, Fry R B, Determination of absolute sound levels and acoustic absorption by thermocouple probes, J Acoust Soc Am 1954;26:294-317) implanted within the tissue. In transparent media, direct observation of the acoustic field can be achieved with optical methods (Raman C V, Nath N S, The diffraction of light by high frequency ultrasonic waves, Proc Indian Acad Sci 1935;2:406-412; Breazeale M A, Heideman E A, Optical methods for measurement of sound pressure in liquids, J Acoust Soc Am 1959;31:24-33) or optical diffraction tomography (Pitts T, Greenleaf J, Lu J Y, Kinnick R, Tomographic schlieren imaging for measurement of beam pressure and intensity, IEEE Ultrasonic Symposium 1994, page 1665). However, these techniques collectively suffer from being either local, invasive, or not applicable to human tissues that are neither transparent nor homogeneous. The ability to provide an accurate non-invasive means, of visualizing the propagation of longitudinal ultrasonic waves in tissue and to quantify its intensity distribution would fill an important need in the development of optimized US therapeutic and imaging strategies.
In current practice, the detection of non-linear US processes is only archived by the use of invasive high bandwidth or tuned hydrophones which measure higher-order harmonics. However, in order to understand the details of shock formation and dissipation of US in heterogeneous tissue, an image of the presence of higher-order motions is needed. While the simultaneous use of multiple hydrophones throughout the media can approximate this distribution it is cumbersome in tissue and will likely alter the US field itself.